NIRS device with optical wavelength and path length correction

ABSTRACT

A near infrared spectrometer and method for wavelength and path length correction are disclosed. The spectrometer includes a number of photodiodes that transmit broadband near infrared measurement light into the tissue and at least one broadband detector which measures the light signal transmitted through the tissue. A processor estimates chromophore concentrations through a comparison of measured light attenuation and modeled light attenuation. The light attenuation model utilizes a light path length distribution derived from a Monte Carlo model and accounts for the spectral shape of the light source as a function of temperature.

CROSS-REFERENCE TO RELATED APPLICATION

This application is a continuation of U.S. patent application Ser. No.13/449,889, filed Apr. 18, 2012, entitled NIRS DEVICE WITH OPTICALWAVELENGTH AND PATH LENGTH CORRECTION, which is incorporated herein byreference in its entirety for all purposes.

TECHNICAL FIELD

The present disclosure relates generally to devices, systems, andmethods for measuring tissue oxygenation in body tissue. Morespecifically, the present disclosure pertains to a near infraredspectrometer and method for optical wavelength and path lengthcorrection.

BACKGROUND

Tissue oxygenation is often used as an indicator of perfusion status inpatients experiencing undifferentiated shock. High risk patients whoreceive continuous monitoring of tissue oxygenation from the trauma baythrough X-ray and CT imaging as well as other procedures have been shownto receive effective interventions sooner, resulting in significantreductions in ICU admission, length of stay, morbidity, and mortality.

Near Infrared Spectroscopy (NIRS) is one technique used fornon-invasively measuring tissue oxygenation. To measure tissueoxygenation, NIRS systems use complex mathematical algorithms thatrelate light attenuation, measured at multiple wavelengths, to theconcentration of different hemoglobin forms, such as oxyhemoglobin(HbO₂) and deoxyhemoglobin (HHb). The concentration of these hemoglobinforms or their calculated oxygen saturation ratio ([HbO₂/[HbO₂+HHb]),defined as saturated oxygen level or “StO₂,” provides an indication ofhow much oxygen is available to the tissue. A clinician may use thesetissue oxygen measurements to evaluate a patient's health status andmake treatment decisions.

NIRS devices employ a light source that illuminates tissue at specificwavelengths of light, typically between 650 nm to 1000 nm, and at leastone photodetector that measures the amount of light exiting the tissuewithin a given area. An example of a NIRS spectrometer is described, forexample, in U.S. Pat. No. 7,239,901, the contents of which areincorporated herein by reference in their entirety for all purposes.

During operation, the amount of light exiting the tissue is compared tothe amount of light emitted into the tissue in order to measure theamount of light lost in the tissue, which is defined as lightattenuation. In tissue, light attenuation occurs from absorption andscattering events. Light absorbing molecules, called chromophores,convert light to heat energy thus reducing the amount of detected light.Light scattering molecules, such as tissue cells and organelles, refractlight thereby changing the direction and hence path length that thelight travels. Although some scattering is required to direct light tothe detector in a reflectance probe configuration, the scattering effecton the light path limits and reduces the amount of light that eventuallyexits the tissue where the photodetector is placed. A reduction indetected light, either from absorption or scattering events, thereforeincreases the amount of light attenuation measured with a NIRSspectrometer.

SUMMARY

The present disclosure pertains to a NIRS spectrometer and method forwavelength and path length correction. A near infrared spectrometer forsensing tissue oxygen measurements in body tissue in accordance with anexemplary embodiment comprises a plurality of light sources configuredto emit broadband, near-infrared measurement light into body tissue; atleast one broadband photodetector configured for sensing at least aportion of the measurement light reflected back from the body tissue; ameans for modeling light attenuations within the body tissue; and ameans for estimating at least one tissue chromophore concentrationwithin the body tissue by comparing attenuations of the sensedmeasurement light reflected back from the body tissue to the modeledlight attenuations.

A NIRS spectrometer for sensing tissue oxygen measurements in bodytissue in accordance with another exemplary embodiment comprises aplurality of light sources configured to emit broadband, near-infraredmeasurement light into body tissue; at least one broadband photodetectorconfigured for sensing at least a portion of the measurement lightreflected back from the body tissue; a temperature sensor configured forsensing a temperature of each light source; a light attenuation modelconfigured for modeling light attenuations within the body tissue basedat least in part on the temperature sensed by the temperature sensor; aprocessor configured for estimating at least one tissue chromophoreconcentration within the body tissue by comparing attenuations of thesensed measurement light reflected back from the body tissue to modeledlight attenuations from the light attenuation model. In someembodiments, the processor is configured to sum the attenuations of themodeled measurement light at a plurality of wavelength increments, thewavelength increments being smaller than a spectral width of each lightsource and a responsivity of the at least one broadband photodetector.

A method for determining one or more tissue oxygen measurements in bodytissue in accordance with an exemplary embodiment comprises coupling aspectrometer to a tissue of interest, the spectrometer including aplurality of light sources configured to emit broadband, near-infraredmeasurement light into the body tissue and at least one broadbandphotodetector configured for sensing at least a portion of themeasurement light reflected back from the body tissue; measuring theattenuation of the measurement light reflected back from the bodytissue; predicting light attenuation within the body tissue using alight attenuation model; and estimating at least one tissue chromophoreconcentration within the body tissue by comparing the attenuation of themeasurement light reflected back from the body tissue to the predictedlight attenuation. In some embodiments, the light attenuation model isconfigured to sum the attenuations of the modeled light at a pluralityof wavelength increments that are smaller than a spectral width of eachlight source and a responsivity of the at least one broadbandphotodetector

While multiple embodiments are disclosed, still other embodiments of thepresent invention will become apparent to those skilled in the art fromthe following detailed description, which shows and describesillustrative embodiments of the invention. Accordingly, the drawings anddetailed description are to be regarded as illustrative in nature andnot restrictive.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a block diagram of a near infrared spectrometer in accordancewith an illustrative embodiment;

FIG. 2 is a perspective view showing several illustrative components ofthe spectrometer of FIG. 1;

FIG. 3 is a schematic view depicting a simulation of light rays orphotons propagating through a scattering medium such as tissue;

FIG. 4 is a flow diagram of a method for determining one or more tissueoxygen measurements in accordance with an illustrative embodiment;

FIG. 5 is a graph showing an example of a modeled path lengthdistribution for a given light source and detector spacing and a giventissue layer;

FIG. 6 is a schematic view of a calibration apparatus in accordance withan illustrative embodiment;

FIG. 7 is a flow diagram of an example process for measuring the powerof a light source as a function of temperature and/or operating voltageusing the apparatus of FIG. 6;

FIG. 8 depicts a number of graphs including several example spectralshape curves for multiple LED light sources, the absorption coefficientof muscle tissue at 70% water and 50% tissue oxygen saturation, thescattering coefficient of muscle tissue, the effective path length, andphotodetector sensitivity;

FIG. 9 is a graph showing the spectral shape curve of an LED at threedifferent temperatures; and

FIG. 10 is a graph comparing the spectral shape curves of an LEDmeasured at 22° C. and predicted at 22° C. after a normalization andinterpolation process has been performed on the spectral response curvesmeasured at 10° C. and 40° C.

While the invention is amenable to various modifications and alternativeforms, specific embodiments have been shown by way of example in thedrawings and are described in detail below. The intention, however, isnot to limit the invention to the particular embodiments described. Onthe contrary, the invention is intended to cover all modifications,equivalents, and alternatives falling within the scope of the inventionas defined by the appended claims.

DETAILED DESCRIPTION

FIG. 1 is a block diagram of a spectrometer 10 in accordance with anillustrative embodiment. As shown in FIG. 1, the spectrometer 10includes a spectrometer interface 12 and a control unit 14. Thespectrometer interface 12 includes a number of light sources and/orlight pathways for directing light into the tissue 16 under study, andfor subsequently collecting measurement light returned back from thetissue 16. In some embodiments, the spectrometer interface 12, controlunit 14, and/or other spectrometer components are integrated into asingle, hand-held device, which eliminates the need for separatefiber-optic cables for transmitting light back and forth between thespectrometer interface 12 and the control unit 14. In one embodiment,the circuitry for the spectrometer interface 12 is located on the sameboard as the circuitry for the control unit 14, and the spectrometer 10is configured to couple directly to the patient's tissue 16 withoutrequiring the attachment of a separate optical probe. In otherembodiments, the light sources are located in the spectrometer interface12, and light pathways are carried in a fiber-optic cable for deliveringand receiving light back and forth between the spectrometer interface 12and a separate optical probe that connects to the patient's tissue 16.

In the embodiment of FIG. 1, the control unit 14 includes an opticaldetector 18 and a controller 20. Light 22 generated by the controller 20is delivered into the patient's tissue 16 via the spectrometer interface12. In some embodiments, a temperature sensor 24 located at or near thelight sources is used to measure the temperature of the light sources,as discussed further herein. Measurement light signals 26 collected bythe optical detector 18 as well as a reference light signal 28 obtainedfrom the tissue, in turn, are transmitted via the spectrometer interface12 to the optical detector 18. Based on these signals 26, 28, theoptical detector 18 produces electrical signals representative of thelight signals at each wavelength of interest. The controller 20 thenprocesses these signals along with the signal from the temperaturesensor 24 to generate data representative of the measured tissueparameter(s), including the saturated oxygen level (StO₂) within thetissue. An example method for determining tissue parameter data such asStO₂ is described further herein with respect to FIG. 4. Othertechniques can also be used for processing the signals and generatingdata representative of tissue oxygenation.

An electrical connector 29 connects the spectrometer 10 to a displayunit 30, which can be used to display the tissue measurement data invarious user-defined formats. In some embodiments, the display unit 30includes an LCD display screen 31 and a user-interface 32. The displayunit 30 can also include other functionality for operating thespectrometer 10, including a power source for supplying power to thespectrometer 10 and/or a computer interface port for connecting thespectrometer 10 to another device for further analysis and/or storage ofthe tissue oxygenation data.

FIG. 2 is a perspective view showing several illustrative components ofthe spectrometer 10 of FIG. 1. FIG. 2 may represent, for example,several illustrative components of the spectrometer interface 12 andcontrol unit 14 of FIG. 1. As shown in FIG. 2, and in some embodiments,the spectrometer 10 includes a spectrometer board 34, a light sourceboard 36, a mixer 38, a first photodetector 40, and a secondphotodetector 42.

The spectrometer board 34 includes circuitry for operating the lightsource board 36 and photodetectors 40, 42, and for converting analogsignals from the photodetectors 40, 42 into corresponding digitalsignals. In some embodiments, the spectrometer board 34 includesprogrammable firmware used to convert these digital signals intoattenuation signals, which as discussed further herein, can be used todetermine an estimated value of saturated tissue oxygenation (StO₂),oxyhemoglobin (HbO₂) and deoxyhemoglobin (HHb), as well as other desiredtissue oxygen parameters. A processor 44 is configured to process themeasurement and reference light signals sensed by the photodetectors anddetermine one or more associated tissue oxygen parameters.

The light source board 36 includes a number of photodiodes 46 (e.g.,light-emitting diodes (LEDs)) that emit near-infrared broadbandmeasurement radiation that is generally centered at wavelengths of 680nm, 720 nm, 760 nm, and 800 nm. The LEDs 46 are each optically coupledto the input end 48 of the mixer 38, and are configured to emitradiation directly into mixer 38 in lieu of first passing the radiationthrough light source conditioning optics prior to entry into the mixer38, as is done by some NIRS spectrometer devices. For example, the LEDs46 are coupled to the input end 48 of the mixer 38 without the use ofinterference filters, focusing lenses, and/or fiber-optic bundlessometimes used to collimate, filter, and direct the emitted radiationfrom each LED source into the mixer 38. The output end 50 of the mixer38, in turn, is optically coupled directly to the tissue to be analyzedwithout being passed through a fiber optic cable, as is also done bysome NIRS spectrometer devices. In some embodiments, film couplers maybe used to prevent instabilities in optical transmission that may becaused by interference effects at the input and output ends 48, 50 ofthe mixer 38. Several example film couplers that are suitable for thispurpose are described further in U.S. Patent Publication No.2011/0102791, the contents of which are incorporated herein by referencein their entirety for all purposes.

From the perspective of spectrometer size, the light source conditioningoptics used in some NIRS spectrometer devices consume a large portion ofthe total spectrometer size, and increase the overall complexity of thedevice. The elimination of the light source conditioning optics in thespectrometer 10 of FIGS. 1-2 reduces the cost and complexity of thespectrometer 10, and allows the LEDs 46 to be positioned in closeproximity together on the light source board 36. For example, thecollimating lenses, interference filters, and focusing lenses used tocollimate, filter, and focus the light in some NIRS devices typicallyrequire precision optical mounts to precisely position these componentsrelative to each other, increasing the cost and complexity of thedevice. Moreover, the elimination of fiber optics on the output end 50of the mixer 38 allows the spectrometer 10 to be directly coupled to thetissue, eliminating the need for a separate optical probe.

The mixer 38 is made from a suitable glass, and is sized and shaped toequally distribute the intensity of the measurement radiation output,thereby creating a spatially uniform illumination at the output 50 froma highly, non-spatially uniform source at the input 48. An example of asuitable glass is Schott SF11® glass, which does not degrade in medicalx-ray environments and which provides desirable optical (e.g.,transmittance, refractive index, dispersion), mechanical, and thermalproperties over the range of wavelengths used by the LEDs 46. In theembodiment shown, the mixer 38 has an essentially constant, rectangularcross-section along its length. Other polygonal cross-sectional shapessuch as, for example, hexagonal and/or octagonal, may also be used. Thecross-sectional shape of the mixer 38 may also vary between the inputand output ends 48, 50.

In some embodiments, the first photodetector 40 is coupled to thespectrometer board 34, and is configured to convert the output lightsignal received from the tissue into an analog electrical signal forprocessing via the processor 44. An ambient light filter 52 isconfigured to filter out any light signal that is not within thespectral range of interest. In some embodiments, for example, the filter52 is configured as a passband filter to filter wavelengths shorter thanapproximately 670 nm and larger than approximately 810 nm. The ambientlight filter 52 may filter out light at different optical passbands,however, depending on the specific wavelength range of the lighttransmitted by the LEDs 46.

In some embodiments, the second photodetector 42 is similarly coupled tothe spectrometer board 34, and is configured to receive a portion of thelight taken through a shorter path of the tissue than the measurementlight detected by the first photodetector 40. In certain embodiments,and as shown in FIG. 2, the first and second photodetectors 40, 42 areeach arranged side-by-side and are positioned at a known, fixed distanceapart from each other.

In use, the coupling of the first and second photodetectors 40, 42directly to the tissue eliminates the need for fiber-optic connectionsbetween the photodetectors 40, 42 and the tissue. The direct coupling ofthe second photodetector 42 to the tissue also eliminates the need for afeedback attenuator, which is used by some NIRS spectrometer devices toreduce the magnitude of the feedback signal to be compatible with thedynamic range of a feedback photodetector. The need for such attenuationtypically arises as a result of coupling the output end of the mixerthrough a feedback fiber optic to the feedback photodetector in lieu ofobtaining the feedback light from the tissue itself, as is done with thespectrometer 10 of FIG. 2.

To compensate for the elimination of the light source condition optics,including the presence of interference filters sometimes used in NIRSspectrometer devices for controlling the spectral properties of theemitted light, a temperature sensor 54 on the light source board 36 canbe used to monitor the temperature of the LEDs 46 during operation. Insome embodiments, the temperature sensor 54 comprises a thermistor orthermocouple that outputs a temperature signal indicative of theoperating temperature of the LEDs 46. Multiple temperature sensors 54can also be used to sense the operating temperature of the LEDs 46.Using the temperature signals, and as discussed further herein, theprocessor 44 can be configured to determine the temperature dependentspectral shape of the LEDs 46 and predict the amount of lightattenuation in the tissue.

When a chromophore exists in a non-scattering medium, the Beer-LambertLaw in equation (1) below defines the absorption of light at onewavelength (A_(λ)) in terms of the measured light intensity entering andexiting the medium (I_(in,λ) and I_(out,λ)):

$\begin{matrix}{A_{\lambda} = {{\ln\left( \frac{I_{{in},\lambda}}{I_{{out},\lambda}} \right)} = {ɛ_{\lambda}{CL}}}} & (1)\end{matrix}$

where:

ε_(λ) is the absorption coefficient at a specific wavelength, definingthe attenuation magnitude per unit path length per unit concentrationfor a specific chromophore;

C is the concentration of the chromophore; and

L is the optical path length representing the linear distance that thelight traveled within the non-scattering medium.

In a scattering medium such as body tissue, however, optical scatteringproduces many different optical path lengths for the light that travelsthrough tissue and reaches the photodetectors 40, 42. These scatteredpath lengths are characterized as being longer than the path lengthsthat would otherwise exist if scattering did not occur. Additionally, intissue, the intensity weighted path length distributions begin to varywith wavelength since absorption magnitude, which is wavelengthdependent via the absorption coefficient (ε_(λ)), increases theprobability that the longer optical path lengths within the distributionof all path lengths are more attenuated before reaching thephotodetectors 40, 42.

An example of this phenomenon in a scattering medium can be seen in FIG.3, which depicts a Monte Carlo (MC) light simulation of three light raysor photons entering into a non-linear scattering medium such as bodytissue. As can be seen in FIG. 3, three example light rays 56, 58, 60are transmitted into tissue 62 having a boundary volume defined byboundary layer 64. The emitted light (I₀) emitted by the light sourceand the received light (I_(RAY)) sensed by the photodetector are at afixed distance apart from each other to permit reflectance modemeasurements to be taken within the tissue 62.

When the light rays 56, 58, 60 enter into the tissue 62, the scatteringcharacteristics of the tissue 62 cause each ray 56, 58, 60 to traverse adifferent optical path within the tissue 62, as shown. The variance inpath lengths of the rays 56, 58, 60 in a scattering medium are theresult of a number of different factors, including the scatteringproperties, geometry, and the boundary optical properties of the medium.

An effective optical path length L_(eff,λ) can be defined generally asan equivalent uniform path length in a non-scattering medium that wouldyield the same attenuation as the varying path length distribution inthe scattering medium. As set forth in equation (2) below, L_(eff,λ) canbe determined by the product of a wavelength dependent differential pathlength factor (a_(λ)) and the optical path length for zero scattering(L):L _(eff,λ) =a _(λ)(L)  (2)

For tissue spectrometers where the light source and photodetector arelocated on the same tissue surface (i.e., reflectance mode positioning),the optical path length for zero scattering (L) in equation (2)represents the separation distance between the light source andphotodetector. Thus as separation distance between the light source andphotodetector increases, the optical path length also increases. If achromophore or combination of chromophores strongly absorbs a particularwavelength of light, then that wavelength of light will have a shortereffective optical path length (L_(eff,λ)).

In addition to the optical path length being different in a scatteringtissue medium, scattering attenuation (A_(scat,λ)) also contributes tototal light attenuation. Scattering attenuation is wavelength dependentbecause in a scattering tissue medium the shorter wavelengths of lightare more effectively scattered than longer wavelengths of light, whichdecreases the probability that the light reaches the photodetector. Thusin body tissue, a modified version of the Beer-Lambert Law isapplicable, as expressed in equation (3) below:

$\begin{matrix}{A_{\lambda} = {{\ln\left( \frac{I_{{in},\lambda}}{I_{{out},\lambda}} \right)} = {{{ɛ_{\lambda}(C)}L_{{eff},\lambda}} + A_{{scat},\lambda}}}} & (3)\end{matrix}$

For a measurement of tissue oxygenation, the detected wavelengths usedin equation (3) above can be selected such that the attenuationmeasurements are sensitive and specific to oxyhemoglobin anddeoxyhemoglobin and less sensitive and specific to background absorbersthat could confound measurement accuracy, such as fat, water, andmelanin. In some cases, these background absorbers may also not requiremeasurement since their concentration levels may have limited or noclinical utility.

In the NIRS wavelength region, both hemoglobin and other backgroundabsorbers have broad overlapping absorption spectra. A more accuratetissue attenuation model can include concentration and absorptivitycoefficient terms for all chromophores that significantly contribute tothe total light attenuation at the measured light wavelengths. Also,since fat and melanin are generally located above muscle tissue, theeffective light path length for these chromophores can be different thanthat for the muscle tissue chromophores; specifically, oxyhemoglobin(HbO₂), deoxyhemoglobin (HHb), oxymyoglobin (MyO₂), deoxymyoglobin (My)and water (H₂O). To account for multiple absorbers and differenteffective light path lengths associated with layered or non-homogenoustissue such as muscle, fat, and skin, the following solution for themodified Beer-Lambert Law is applicable:

$\begin{matrix}{A_{\lambda} = {{\left( {{ɛ_{{HbO}\; 2\lambda}C_{{HbO}\; 2}} + {ɛ_{{Hb}\;\lambda}C_{Hb}} + {ɛ_{{MyO}\; 2\lambda}C_{{MyO}\; 2}} + {ɛ_{{My}\;\lambda}C_{My}} + {ɛ_{H\; 2O\;\lambda}C_{H\; 2O}}} \right){Leff}_{{muscle},\lambda}} + {\left( {{ɛ_{{melanin}\;\lambda}C_{melanin}} + {ɛ_{{HbO}\; 2\lambda}C_{{HbO}\; 2}} + {ɛ_{{Hb}\;\lambda}C_{Hb}} + {ɛ_{H\; 2O\;\lambda}C_{H\; 2O}}} \right){Leff}_{{skin},\lambda}} + {\left( {{ɛ_{{fat}\;\lambda}C_{fat}} + {ɛ_{{HbO}\; 2\lambda}C_{{HbO}\; 2}} + {ɛ_{{Hb}\;\lambda}C_{Hb}} + {ɛ_{H\; 2O\;\lambda}C_{H\; 2O}}} \right){Leff}_{{fat},\lambda}} + A_{{{muscle}\mspace{14mu}{scat}},\lambda} + A_{{{skin}\mspace{14mu}{scat}},\lambda} + A_{{{fat}\mspace{14mu}{scat}},\lambda}}} & (4)\end{matrix}$

In order to compensate for the unknown variables in equation (4) above,the spectrometer 10 is configured to measure light attenuation atseveral different wavelengths. These unknown variables include all thechromophore concentrations, the effective light path lengths, andscattering attenuations. The absorptivity coefficients are known, andcan be preprogrammed within the spectrometer. A simultaneous algebraicsolution of equation (4) in which the number of wavelength specificattenuation measurements equals or exceeds the number of unknowns ispossible, but has limitations. One limitation is that the absorbance ofeach chromophore may all increase or decrease together resulting in aco-linearity that causes the equation solution to become unstable as thenumber of wavelengths increases. Also, it is possible that differentcombinations of the unknown variables produce a similar solutionresulting in an erroneous tissue oxygenation measurement. Anotherlimitation of solving equation (4) for all the unknowns is related tothe complexity and cost of the optical hardware needed to produce andmeasure the numerous wavelengths of light.

The number of unknown variables in equation (4) can be minimized so thatthe number of light sources, and hence attenuation measurements, can beminimized. If, for example, the tissue measurement site has skin, fat,and muscle layers, two photodetectors at different separation distancesfrom the light sources can be used. If the attenuation measurements atthe short (i.e., shallow) spacing provide a measurement depth containedwithin the skin and fat layers and the longer (i.e., deep) separationdistance photodiode measures attenuation from skin, fat, and muscle,then subtracting the short distance attenuation from the simultaneouslymeasured longer distance attenuation provides suppression or cancelingof the skin and fat contributions. In this case, the skin and fatvariables of equation (4) can be ignored and removed from the solution.

Another simplification of equation (4) can be made when the measurementtissue is mostly muscle and contains a limited number of chromophoresthat contribute to the overall attenuation. For example, the thenareminence of the hand's palm, which is the muscle between lower thumbjoint and the wrist, does not grow fat tissue like other areas of thebody. Also, skin pigment or melanin is less variable on the palmsurface. Furthermore, for muscle tissue, myoglobin and hemoglobin have anearly indistinguishable profile of absorptivity versus wavelength, andboth chromophores transport oxygen to tissue. Therefore, a commonassumption for NIRS muscle measurements is that the tissue oxygenationrepresents combined hemoglobin and myoglobin effects with a majority ofthe signal being derived from hemoglobin. For this simplification, theHbO₂ and HHb abbreviations represent both hemoglobin and myoglobin.Wider separation distances between a light source and photodetectorweight the optical path length and absorption to the deeper tissuedepths that generally contain much more hemoglobin concentration thanskin. Thus for large light source and photodetector separationdistances, such as 15 mm or more, the skin variables of equation (4) maybe ignored and removed from the solution.

Wavelength dependent scatter attenuation and chromophores other thanhemoglobin/myoglobin and water generally have a constant slope orflatness within the NIRS wavelength region. For wavelengths that span aregion in which the hemoglobin absorptivity coefficient is highlynonlinear with respect to wavelength, such as 680 to 800 nm, then thescattering attenuation and other chromophore attenuation contributionscan be represented by a wavelength dependent slope and offset. For thissimplification, the tissue's scattering attenuation and linearbackground absorbers can be represented as an unknown slope (m) andunknown offset (b) for a linear equation that relates wavelength (λ) toa tissue's background attenuation and absorption (A_(λ)) based on thefollowing expression:A _(λ) =mλ+b  (5)

Equations (3) to (5) can be arranged and combined to form a simplifiedequation in which the optical path lengths are predominately from musclemeasurements, such as when using a 15 mm light source and photodetectorseparation located over the thenar eminence muscle:

$\begin{matrix}{A_{\lambda} = {\ln\left( \frac{I_{{in},\lambda}}{I_{{in},\lambda}\exp^{({{{({{ɛ_{{HbO}\; 2\lambda}C_{{HbO}\; 2}} + {ɛ_{{Hb}\;\lambda}C_{Hb}} + {ɛ_{H\; 2O\;\lambda}C_{H\; 2O}}})}{Leff}_{{tissue},\lambda}} + {m\;\lambda} + b})}} \right)}} & (6)\end{matrix}$

In equation (6) above, the right side of the equation is defined interms of the incoming light intensity, allowing a wavelengthcompensation and correction integration calculation to be made thatrelies on pre-characterizing and predicting the wavelength and intensitycharacteristics of each light source, as set forth in equation (7)below:

$\begin{matrix}{A_{LED} = {\ln\left( \frac{\sum\limits_{\lambda\mspace{14mu}\min}^{\lambda\mspace{14mu}\max}I_{{in},\lambda}}{\sum\limits_{\lambda\mspace{14mu}\min}^{\lambda\mspace{14mu}\max}{I_{{in},\lambda}\exp^{({{{(\mu_{a,\lambda})}{Leff}_{{tissue},\lambda}} + {m\;\lambda} + b})}}} \right)}} & (7)\end{matrix}$

where:

μ_(a,λ)=ε_(HbO2λ)C_(HbO2)+ε_(Hbλ)C_(Hb)+ε_(H2Oλ)C_(H2O) is the tissues'wavelength dependent absorption coefficient represented as the summedproduct of each chromophore's absorptivity coefficient andconcentration.

Equation (7) above relates attenuation measured with one light sourceand photodetector to the individual concentrations of significant,non-linear absorbing chromophores such as hemoglobin and water, and acomposite of linear attenuators such as melanin or bilirubin. Includingthe numeric integration terms over the distribution of wavelengths(λ_(min) to max) contained within a light source appropriately weightsthe wavelength dependent variables to the exact wavelengths emitted intothe tissue, thus enabling improved accuracy of the desired concentrationmeasurements. Since thenar eminence tissue is robust to edema, the waterconcentration can be assumed constant, 70 wt %. Also, since effectivepath length for each wavelength can be predicted from additionalmodeling and pre-characterization, the unknowns of equation (7) reducesto four. As such, four light sources 46 each with different mean orcentral wavelengths are used by the spectrometer 10 to solve equation(7) and measure tissue hemoglobin concentrations and oxygenation levels.

FIG. 4 is a flow diagram of an illustrative method 66 for determiningone or more tissue oxygen measurements in body tissue such as saturationoxygen level (StO₂) and/or chromophore concentration. The method 66 mayrepresent, for example, an algorithm or routine run by the spectrometer10 of FIGS. 1-2 to provide wavelength compensation and path lengthcorrections needed to measure tissue oxygenation when the opticalhardware of the spectrometer 10 does not control or precisely define thewavelengths of light emitted by the LEDs 46. In one embodiment, forexample, the method 66 can comprise an algorithm or routine run by theprocessor 44 to compensate for wavelength and path length effects thatresult from the elimination of the light source conditioning optics usedby some NIRS spectrometer devices.

The method 66 may begin generally at block 68, in which the operatingtemperature or voltage of a multiple wavelength light source ismeasured. With respect to the spectrometer of FIGS. 1-2, and in someembodiments, the operating temperature can be determined using thetemperature sensor 54 coupled to the light source board 36. In otherembodiments, the operating voltage can be sensed in lieu of temperature,which can then be used to derive the operating temperature of the LEDs46.

The wavelength dependent intensity variable of equation (7) can bepre-characterized and calibrated using a calibration source (block 70).The calibration process can include, for example, obtaining a photodiodespectral shape reading from one or more of the LEDs 46 at one or moretemperatures that can be used to correct for any temperature effect onthe output response of equation (7). An example apparatus and processthat can be used for pre-characterizing and calibrating thespectrometer's 10 light source and photodetector are described furtherherein with respect to FIGS. 6 and 7, respectively.

The spectrometer 10 can be coupled to a tissue of interest containingchromophore samples to be measured (block 72), which can be sensed viathe first photodetector 40 and using the temperature and wavelengthdependent light intensity or power values obtained from the calibrationapparatus. At block 74, the tissue optical attenuation can then beempirically measured by at least one LED 46 and one photodetector 40 ofthe spectrometer 10.

In some embodiments, the spectrometer algorithm is configured toestimate or guess the amounts of tissue chromophores contained in thetissue of interest (block 76). In certain embodiments, for example, thealgorithm may utilize equation (7) discussed herein to relateattenuation measured with the at least one LED and photodetector to theindividual concentrations of significant, non-linear absorbingchromophores such as hemoglobin and water, and a composite of linearattenuators such as melanin, bilirubin, and scattering attenuation.Other means for estimating or guessing the amounts of tissuechromophores present in the tissue can also be employed.

In some embodiments, the spectrometer 10 may estimate the effectiveoptical path length (L_(eff,λ)) of the traversed light within the tissue(block 78). To estimate L_(eff,λ) for the wavelength distribution (i.e.,spectral shape) of each photodiode light source, the algorithm mayreceive, as inputs, the estimated chromophore concentrations (block 76)and the known chromophore absorptive properties (block 80). Thealgorithm may also receive, as an input, the path length distribution(block 80) obtained by applying the known tissue light scatteringproperties and geometry of the tissue to be measured (block 82) to aMonte Carlo simulation model or the like (block 84). From theseparameters, the spectrometer algorithm estimates the effective opticalpath length (L_(eff,λ)), and provides this estimate to a tissue lightattenuation model (block 86). Since the wavelength characteristics ofthe LED are not precisely defined due to the elimination of light sourceconditioning optics, at block 88 the spectrometer algorithm also inputsestimates of the wavelength dependent power properties of the LED fromthe measured LED operating temperature sensed by the temperature sensorat block 68, and provides this as an input to the tissue lightattenuation model. In some embodiments, the photodetector sensitivitycan also be measured (block 90) and provided as an input to the tissuelight attenuation model.

Based on the estimate of the effective optical path length (L_(eff,λ))and the predicted light source power properties, the tissue lightattenuation model predicts the light attenuation within the tissue(block 92). In certain embodiments, for example, the spectrometer 10 maypredict the tissue light attenuation by solving for the lightattenuation (A_(LED)) using the summation of discrete attenuationcalculations (the numerical integration calculation expressed inequation (7)) over a distribution of wavelengths (λ_(min) to λ_(max))defined by the spectral shape of the LED. Thus, a unique light sourceintensity, path length, and absorptivity coefficient are input into theattenuation model for each wavelength increment defined by each lightsource's wavelength distribution or spectral shape. In some embodiments,the wavelength increments are smaller than the spectral width of eachlight source and the responsivity of the photodetector. An example of anon-linear optimization algorithm or routine that can be used forsolving equation (7) is described further herein with respect to FIG. 7.

At decision block 94, the predicted tissue light attenuation is thencompared against the measured tissue light attenuation to assess if theattenuation value matches or is within a useful degree of accuracy(e.g., ≤2% absolute difference). In some embodiments, for example, amatch may be determined by performing a root mean square function on thedata and determining whether a root mean square error is minimized,indicating the likelihood of a match. If no match is found, then newestimates for the chromophore amounts are selected by the spectrometer10 (block 96) and the process of estimating an effective path lengthbased on the known chromophore absorption and tissue light scatteringproperties and the predicted light source power properties is applied tothe tissue light attenuation model to generate another predicted valueof the tissue light attenuation. In some embodiments, this process is aniterative solving process that is repeated multiple times until thepredicted attenuation values match or are sufficiently close to themeasured tissue light attenuation values.

If, on the other hand, the spectrometer 10 determines that the predictedtissue light attenuation matches or is within a useful degree ofaccuracy, the spectrometer 10 may display and trend the correctlypredicted chromophore concentrations (block 98). For example, and insome embodiments, the spectrometer 10 may display the tissue chromophoreamounts on a display screen coupled to the spectrometer 10. The tissuechromophore amounts can also be converted into other forms such as StO₂,allowing the physician to quickly determine, in real-time, the StO₂level within the tissue. To determine and display an StO₂ value, forexample, the spectrometer 10 would perform the method 66 for twochromophores relating to oxy and deoxy hemoglobin. Historical data takenover a period of time may also be displayed on the display screen,providing the physician with trending data related to changes in thepatient's health status over time.

FIG. 5 is an example graph 100 showing path length distribution ofsample light rays 102 through body tissue obtained by Monte Carlosimulation modeling. FIG. 5 may represent, for example, a normalizedgraph 100 of path length distribution of light rays 102 in body tissuein which the absorption coefficient (μ_(a)) is near zero, the scatteringcoefficient (μ_(s)′) is 8 cm⁻¹, the effective optical path length(L_(eff,λ)) is 19 cm⁻¹, and the light source to photodetector spacing is15 mm.

To simulate light traveling in tissue, the Monte Carlo simulation modelcan be used by the spectrometer to determine and record the position ofa light ray after each scattering event. The Monte Carlo model can beused, for example, for predicting light energy penetration within atissue such as in radiation or photodynamic therapy where the optimallight energy or dose is dependent upon knowing a tissue's opticalproperties. Also, if tissue light attenuation is known, then the MonteCarlo model can be used to measure tissue optical properties, such asabsorption or scattering, to help identify tissue pathologies such ascancer. In some embodiments, the Monte Carlo model can be used to obtaina distribution of light intensities or flux striking the photodetectorfor an anticipated range of scattering conditions with a fixedabsorption coefficient that is at or near zero, and based on a fixedspacing between the light source and photodetector. In some embodiments,the simulation can be performed by a software program, firmware, and/orhardware. In one embodiment, for example, the simulation can beperformed by a software program such as TRACEPRO® available from LambdaResearch Corporation of Littleton, Mass.

The Monte Carlo model requires a number of input variables, includinglight source variables such as location, wavelength, illumination area,and angular distribution; detector variables such as viewing angle,detector area, and location; and tissue boundary and optical propertiessuch as width, depth, shape, refractive index, absorption, scattering,and anisotropy. Based on these input variables, the Monte Carlo modellaunches light photons into the simulated tissue and then determines andtracks the stepwise movement of the photons that reach thephotodetector. The photon's step size is randomly sampled from aprobability distribution for the free path between tissue interactionevents and launched into the tissue. The inverse value of the mean freepath equals the transport scattering coefficient. At the end of a photonstep movement, the photon number or intensity is attenuated according tothe absorption properties per unit path length (i.e., the absorptioncoefficient).

After absorption, a new photon direction is randomly chosen from aprobability distribution for the deflection angle of the scattered lightusing the tissue's anisotropy coefficient. This process can be repeatedone or more times until either the photon reaches the photodetector oris eventually lost such as being fully absorbed or exiting the tissueboundary. The software program then returns the intensity of each lightray that exits the tissue where the photodetector is located.

The light intensity distribution for the numerous photon or raysimulations is then used in combination with the Beer-Lambert Law tocalculate the effective optical path length. From the Beer-Lambert Law,and in some embodiments, the effective optical path length of eachdetected ray (L_(RAY)) can be calculated based on the followingequation:L _(RAY)=Log(I ₀ /I _(RAY))/μ_(a)  (8)

where I₀=1 and (μ_(a)) is the absorption coefficient.

A constant absorption coefficient (μ_(a)) such as 0.01 can be used inequation (8) above so that the Monte Carlo simulation is performed onlyfor the differing scattering properties of tissue, thereby significantlydecreasing the number of required simulations. This can be accomplishedbecause equation (8) can be used to define the effective path lengths atthe differing absorption coefficient values.

From the Beer-Lambert Law, the intensity of each detected light ray(L_(RAY)) may be calculated for numerous absorption coefficient values(μ_(a)) that are not zero, and which span the range of probable tissuevalues. By way of example and not limitation, L_(RAY) may be solved for43 pa values ranging from 0.001 to 11.7 using the following equation:I _(RAY) =I ₀exp−(μ_(a)(L _(RAY)))  (9)

The effective optical path length (L_(eff,λ)) can then be calculated foreach of the μ_(a) conditions from equation (9) above as follows:

$\begin{matrix}{L_{{eff},\lambda} = \frac{{Ln}\left( {\sum{I_{0}/{\sum I_{Ray}}}} \right)}{\mu_{a,\lambda}}} & (10)\end{matrix}$

The process can be repeated for the simulated media having scatteringproperties resembling tissue at different wavelengths. The absorptionand transport scattering coefficients μ_(a) and μ_(s)′ have units ofreciprocal path length (cm⁻¹), and for tissue typically range from 0cm⁻¹ to 0.5 cm⁻¹ and 5 to 10 cm⁻¹, respectively. As an example, theprocess may be repeated for different wavelengths at μ_(s)′ of 5, 8, and10 cm⁻¹.

The effective optical path length (L_(eff,λ)) can be obtained for avariety of optical absorption (μ_(a)) and scattering coefficient(μ_(s)′) properties. Since μ_(a) equals the product of the chromophoreconcentration (C) and chromophore absorption coefficient (ε), theestimated chromophore concentration and known absorption coefficientthat are inputted into the tissue attenuation model can be used toselect an appropriate L_(eff,λ) for a given wavelength of light. Thewavelength of the light also defines the appropriate μs′ when selectingL_(eff,λ).

In some embodiments, the Monte Carlo model is configured to generate atwo dimensional lookup table that outputs the effective optical pathlength for the following input values of absorption coefficient(μ_(a,λ)) and transport scattering coefficient (μ′_(s,λ)) for a givenwavelength (λ):

$\begin{matrix}{{\mu_{a,\lambda} = {\sum\limits_{i}\left( {ɛ_{i,\lambda}C_{i}} \right)}};} & (11)\end{matrix}$

where i=each measured chromophore concentration; andμ′_(s,λ)=μ_(s,λ)(1−g);  (12)

where μ_(s,λ) is the scattering coefficient and g is the anisotropycoefficient of the tissue. Tissue anisotropy coefficient (g) values aretypically near 0.90, indicating that tissue scattering occurs mostly inthe forward direction.

Other techniques for modeling the relationship of the effective opticalpath length versus the absorption and scatter coefficients can also beemployed. In one alternative embodiment, for example, a Monte Carlotechnique can be performed on thousands of rays for several scatteringconditions in order to directly obtain a distribution of optical pathlengths (e.g., L_(RAY) of equation (8)) without having to model anddefine the intensity distribution of each ray. In another embodiment, atime-resolved system can be used to empirically measure light pathlength distribution using tissues or tissue phantoms having controlledand known scattering profiles.

FIG. 6 is a schematic view of an example calibration apparatus 116 forcalibrating the spectrometer 10 of FIG. 2 in order to predict thewavelength dependent power output of an LED source 46 as a function oftemperature and/or operating voltage. In some embodiments, for example,the calibration apparatus 116 can also be used for calibrating thespectrometer photodetector as part of the calibration step (block 70)described with respect to FIG. 4.

As shown in FIG. 6, the calibration apparatus 116 includes amonochromator 118, an integrating sphere 120 coupled to a spectrometer10 including at least one light source and photodetector to becharacterized, and a reference spectrometer 122 that is substantiallylinear over a wide dynamic range. If the photodetector calibration isnot required, such as when the photodetector's spectral response is notsignificantly different over the wavelength region of interest, then themonochromator 118 would not be needed.

The monochromator 118 serves as a wavelength calibration source, and isconfigured to transmit a sufficiently narrow wavelength band of lightinto the integrating sphere 120. In some embodiments, for example, themonochromator 118 is configured to transmit light having a bandwidth ator less than about 1 nm. The integrating sphere 120 includes a hollowcavity 124 that is coated for high diffuse reflectivity, and serves as adiffuser to provide uniform scattering of the light rays received by thetissue oxygenation and reference spectrometers 10, 122.

FIG. 7 is a flow diagram of an example process 126 for predicting thewavelength specific power of an LED light source as a function oftemperature and/or operating voltage using the calibration apparatus 116of FIG. 6. FIG. 7 may represent, for example, several illustrative stepsthat can be used for determining the light source wavelength dependentpower properties of each LED 46 used by the spectrometer 10 of FIG. 2.

The process 126 can begin generally at block 128, in which the LED to becharacterized is turned off, and the monochromator 118 is activated tosweep a wavelength region of interest. At each step during the sweep,the apparatus 116 records the monochromator wavelength and power readingfrom the reference spectrometer 122 (e.g., as watts), and the relativeintensity of the tissue oxygenation photodetector as digital counts. Theoutput of this step provides a photodetector spectral shape curve(S_(D,λ)) in counts per watt for each photodetector, as shown, forexample, in the bottom curve of FIG. 8. In another step, themonochromator 118 is turned off and the LED 46 to be characterized forthe tissue oxygenation spectrometer 10 is turned on (block 132). Thereference spectrometer 122 records the intensity versus wavelength datafor the LED at a certain temperature (block 134). The intensity andwavelength values of this step may provide an LED spectral shape curve(P_(LED,λ)) such as that shown, for example, in FIGS. 8 and 9.

FIG. 8 depicts a number of graphs including several example spectralshape curves 136 a-136 for multiple LED light sources (P_(LED)), theabsorption coefficient of muscle tissue at 70% water and 50% tissueoxygen saturation 138 (“μ_(a)′”), the scattering coefficient of muscletissue 140 (“μ_(s)′”), the effective path length 142 (“L_(eff)”), andphotodetector sensitivity 144 (“S_(D)”). The output spectral shape curvefrom the characterization process 126 of FIG. 7 can be used inconjunction with the absorptivity and scattering propertiescorresponding to each wavelength increment (as indicated by the dashedvertical lines in FIG. 8) to predict the LED signal attenuation(A_(LED)) at a particular temperature that is numerically solved overthe entire wavelength range of each LED employed by the spectrometer 10,as indicated in equation (13) below:

$\begin{matrix}{A_{LED} = {\ln\left( \frac{\sum\limits_{\lambda\mspace{14mu}\min}^{\lambda\mspace{14mu}\max}{P_{{LED},\lambda}S_{D,\lambda}}}{\sum\limits_{\lambda\mspace{14mu}\min}^{\lambda\mspace{14mu}\max}{P_{{LED},\lambda}S_{D\;\lambda}\exp^{({{{(\mu_{a,\lambda})}{Leff}_{{tissue},\lambda}} + {m\;\lambda} + b})}}} \right)}} & (13)\end{matrix}$

Any scaling LED power or photodetector sensitivity scalar variable thatis not wavelength specific is factored out of the numerically solvedintegral in both the numerator and denominator portion of equation (13),and thus would cancel out.

The characterization process 126 can then be repeated for multipledifferent temperatures in order to determine the relative intensity andwavelength values of the LED (P_(LED,λ)) across a temperature spectrum.In some embodiments, for example, the calibration apparatus 116 isconfigured to record this information over a large number oftemperatures (e.g., one value every 1° C.), and use the closest databased on the measured temperature.

In another embodiment, the information is recorded over only a fewtemperatures (e.g., the low and high temperature extremes), and the datais interpolated to obtain a result in between. To correctly compensatefor the LED spectral shape, each point along the spectral distributionshould be interpolated or extrapolated rather than just being shiftedbased on the peak or centroid of the response curve. This is due to thenon-linear temperature characteristics of the LED. For example, as theLED is heated, the LED experiences not only a shift towards a longerwavelength, but the distribution is also broadened and the peakintensity of the light is reduced. Conversely, as the LED is cooled, theLED experiences a shift towards a shorter wavelength, a narrowing of thedistribution, and a larger peak intensity.

To compensate for these LED spectral shape distortions at differenttemperatures, a normalization process can be performed on the spectralshapes for each of the “hot” and “cold” temperature values. In someembodiments, for example, the “hot” and “cold” spectral shapes can befirst normalized based on their peak intensity, and the intensity andwavelength axes can be transposed prior to interpolation to predict theLED spectral shape at any given LED operating temperature.

FIGS. 9-10 are several graphs showing the spectral shape of opticalintensity versus wavelength for an LED at two different temperatures,representing the lower and upper bounds of the operating temperature ofthe LED. In a first graph 146 shown in FIG. 9 prior to any compensation,the spectral response curve 148 for the LED operating at a relativelycold temperature (e.g., at 10° C.) is shifted to the left of the roomtemperature response curve 150 of the LED (e.g., at 22° C.). A secondspectral response curve 152 associated with the LED operating at arelatively hot temperature (e.g., 40° C.), in turn, is shifted to theright of the actual response curve 150. In addition, and as can befurther seen in FIG. 9, the shape of the spectral response curve 148 forthe relatively cold LED is narrower and has a larger peak opticalintensity than that of the spectral response curve 152 for therelatively hot LED, which has a broader shape with a smaller peakoptical intensity.

FIG. 10 is a graph 154 showing the spectral shape curves of the LEDsafter a normalization and interpolation process has been performed onthe peak intensity and wavelength values. As shown in FIG. 10, whennormalized spectral response curves 156, 158 for the 10° C. and 40° C.LEDs are generated, an interpolation of the two curves 156, 158 along aline of uniform intensity generates a predicted spectral response curve160 at 22° C. that closely matches the actual measured spectral shapecurve 162 of the LED at 22° C.

Various modifications and additions can be made to the exemplaryembodiments discussed without departing from the scope of the presentinvention. For example, while the embodiments described above refer toparticular features, the scope of this invention also includesembodiments having different combinations of features and embodimentsthat do not include all of the above described features.

What is claimed is:
 1. A near infrared spectrometer for sensing tissueoxygen measurements in body tissue, comprising: a plurality of lightsources configured to emit broadband, near infrared measurement lightinto body tissue; at least one broadband photodetector configured forsensing at least a portion of the measurement light reflected back fromthe body tissue; a means for modeling light attenuations within the bodytissue, the means including a light attenuation model configured to sumpredicted attenuations of the light in tissue at a plurality ofwavelength increments, wherein the wavelength increments are smallerthan a spectral width of each light source and a responsivity of the atleast one broadband photodetector; and a means for estimating at leastone tissue chromophore concentration within the body tissue by comparingattenuations of the sensed measurement light reflected back from thebody tissue to the modeled light attenuations.
 2. The spectrometer ofclaim 1, further comprising a temperature sensor configured for sensinga temperature of the plurality of light sources.
 3. The spectrometer ofclaim 2, wherein the light attenuation model is configured to compensatefor a spectral shape of each light source based at least in part on thetemperature.
 4. The spectrometer of claim 3, wherein the spectral shapeof each light source is pre-characterized at a plurality oftemperatures.
 5. The spectrometer of claim 4, wherein the spectral shapeof each light source is interpolated or extrapolated from a plurality ofpre-characterized temperatures.
 6. The spectrometer of claim 1, whereina path length distribution for each wavelength increment of themeasurement light is predicted using a simulation model.
 7. Thespectrometer of claim 6, wherein the simulation model is a Monte Carlomodel.
 8. The spectrometer of claim 1, wherein the plurality of lightsources are each coupled to a light mixer having an input end and anoutput end.
 9. The spectrometer of claim 8, wherein the output end ofthe mixer is configured to couple directly to the tissue.
 10. Thespectrometer of claim 8, wherein the at least one broadbandphotodetector is configured to couple directly to the tissue.
 11. Thespectrometer of claim 1, further comprising a secondary photodetectorconfigured for sensing the measurement light reflected back from thetissue.
 12. The spectrometer of claim 11, wherein the secondaryphotodetector is configured for sensing at least a portion of themeasurement light having a shorter optical path than an optical path ofthe measurement light sensed by the at least one broadbandphotodetector.
 13. The spectrometer of claim 12, wherein the at leastone broadband photodetector and secondary photodetector are coupleddirectly to the tissue and are configured to obtain reflectance modemeasurements of the measurement light within the tissue.
 14. A nearinfrared spectrometer for sensing tissue oxygen measurements in bodytissue, comprising: a plurality of light sources configured to emitbroadband, near infrared measurement light into body tissue; at leastone broadband photodetector configured for sensing at least a portion ofthe measurement light reflected back from the body tissue; at least onetemperature sensor configured for sensing a temperature of the pluralityof light sources; a processor configured for modeling light attenuationswithin the body tissue based at least in part on the temperature sensedby the temperature sensor using a light attenuation model and estimatingat least one tissue chromophore concentration within the body tissue bycomparing attenuations of the sensed measurement light reflected backfrom the body tissue to modeled light attenuations from the lightattenuation model; and wherein the processor is configured to sum theattenuations of the light in tissue at a plurality of wavelengthincrements, the wavelength increments being smaller than a spectralwidth of each light source and a responsivity of the at least onebroadband photodetector.
 15. A method for determining one or more tissueoxygen measurements in body tissue, the method comprising: coupling aspectrometer to a tissue of interest, the spectrometer including aplurality of light sources configured to emit broadband, near infraredmeasurement light into the body tissue and at least one broadbandphotodetector configured for sensing at least a portion of themeasurement light reflected back from the body tissue; measuring theattenuation of the measurement light reflected back from the bodytissue; predicting light attenuation within the body tissue using alight attenuation model, wherein the light attenuation model isconfigured to sum the predicted attenuations of light in tissue at aplurality of wavelength increments, the wavelength increments beingsmaller than a spectral width of each light source and a responsivity ofthe at least one broadband photodetector; and estimating at least onetissue chromophore concentration within the body tissue by comparing theattenuation of the measurement light reflected back from the body tissueto the predicted light attenuation.
 16. The method of claim 15, furthercomprising sensing a temperature of each light source, and wherein thelight attenuation model is configured to compensate for a spectral shapeof each light source based at least in part on the temperature.
 17. Themethod of claim 16, wherein predicting light attenuations within thebody includes pre-characterizing the spectral shape of each light sourceat a plurality of temperatures.
 18. The method of claim 16, whereinpredicting light attenuations within the body includes normalizing anintensity distribution based on a peak intensity and interpolating orextrapolating a spectral shape of each light source along a line ofuniform intensity.
 19. The method of claim 16, wherein predicting lightattenuations within the body tissue using a light attenuation modelincludes simulating the path lengths of measurement light within thetissue using a simulation model.